Cardiac output, the volumetric rate at which blood is pumped through the heart, is most often determined clinically by injecting a bolus of chilled saline or glucose solution into the right auricle or right ventricle through a catheter. A thermistor disposed in the pulmonary artery is used to determine a temperature-time washout curve as the chilled injectate/blood mixture is pumped from the heart. The area under this curve provides an indication of cardiac output. Although this thermo-dilution method can give an indication of cardiac output at the time the procedure is performed, it cannot be used for continuously monitoring cardiac output. Moreover, the frequency with which the procedure is performed is limited by its adverse effects on a patient, including the dilution of the patient's blood that occurs each time the chilled fluid is injected. In addition, the procedure poses an infection hazard to medical staff from blood contact, and to the patient, from exposure to possibly contaminated injectate fluid or syringes.
Alternatively, blood in the heart can be chilled or heated in an injectateless method by a heat transfer process using a temperature-conditioned fluid that is pumped in a closed loop, toward the heart through one lumen within the catheter and back through another lumen. The principal advantages of using such a non-injectate heat transfer process to change the temperature of blood are that the blood is not diluted, and the temperature differential between the blood and the heat exchanger is much less compared to the temperature differential between an injectate fluid and blood in the typical thermo-dilution procedure.
U.S. Pat. No. 4,819,655 (Webler) discloses an injectateless method and apparatus for determining cardiac output. In Webler's preferred embodiment, a saline solution is chilled by a refrigeration system or ice bath and introduced into a catheter that has been inserted through a patient's cardiovascular system into the heart. The catheter extends through the right auricle and right ventricle and its distal end is disposed just outside the heart in the pulmonary artery. A pump forces the chilled saline solution through a closed loop fluid path defined by two lumens in the catheter, so that heat transfer occurs between the solution and blood within the heart through the walls of the catheter. A thermistor disposed at the distal end of the catheter monitors the temperature of blood leaving the heart, both before the chilled fluid is circulated through the catheter to define a baseline temperature, and after the temperature change in the blood due to heat transfer with the chilled saline solution has stabilized. Temperature sensors are also provided to monitor both the temperature of the chilled saline solution at or near the point where it enters the catheter (outside the patient's body) and the temperature of the fluid returning from the heart. In addition, the rate at which the chilled solution flows through the catheter is either measured or controlled to maintain it at a constant value. Cardiac output (CO) is then determined from the following equation: ##EQU1## where V.sub.I equals the rate at which the chilled fluid is circulated through the catheter; .DELTA.T.sub.I equals the difference between the temperature of the chilled fluid input to the catheter and the temperature of the fluid returning from the heart; .DELTA.T.sub.B equals the difference between the temperature of the blood leaving the heart before the chilled fluid is circulated and the temperature of the blood leaving the heart after the chilled fluid is circulated (after the temperature stabilizes); and C is a constant dependent upon the blood and fluid properties. The patent also teaches that the fluid may instead be heated so that it transfers heat to the blood flowing through the heart rather than chilled to absorb heat from the blood.
U.S. Pat. No. 4,819,655 further teaches that the cardiac monitoring system induces temperature variations in the pulmonary artery that are related to the patient's respiratory cycle and are therefore periodic at the respiratory rate. Accordingly, Webler suggests that the signal indicative of T.sub.B ' (the temperature of the chilled blood exiting the heart) should be processed through a Fourier transform to yield a period and amplitude for the respiratory cycle, the period or multiples of it then being used as the interval over which to process the data to determine cardiac output.
Instead of cooling (or heating) the blood in the heart by heat transfer with a circulating fluid, the blood can be heated with an electrical resistance heater that is disposed on a catheter inserted into the heart as taught by H. H. Khalil in U.S. Pat. No. 3,359,974. The apparatus required for this type of injectateless cardiac output measurement is significantly less complex than that required for circulating a fluid through the catheter. An electrical current is applied to the resistor through leads in the catheter, and the current is adjusted to develop sufficient power dissipation to produce a desired temperature rise signal in the blood. However, care must be taken to avoid using a high power that might damage the blood by overheating it. An adequate signal-to-noise ratio is instead preferably obtained by applying the electrical current to the heater at a frequency corresponding to that of the minimum noise generated in the circulatory system, i.e., in the range of 0.02 through 0.15 Hz. U.S. Pat. No. 4,236,527 (Newbower et al.) also describes such a system, and more importantly, describes a technique for processing the signals developed by the system to compensate for the effect of the mixing volume in the heart and cardiovascular system of a patient, even one with a relatively large heart. (Also see J. H. Philip, M. C. Long, M. D. Quinn, and R. S. Newbower, "Continuous Thermal Measurement of Cardiac Output," IEEE Transactions on Biomedical Engineering, Vol. BMI 31, No. 5, May 1984.)
Newbower et al. teaches modulating the thermal energy added to the blood at two frequencies, e.g., a fundamental frequency and its harmonic, or with a square wave signal. Preferably, the fundamental frequency equals that of the minimal noise in the cardiac system. The temperature of the blood exiting the heart is monitored, producing an output signal that is filtered at the fundamental frequency to yield conventional cardiac output information. The other modulation frequency is similarly monitored and filtered at the harmonic frequency, and is used to determine a second variable affecting the transfer function between the injection of energy into the blood and the temperature of the blood in the pulmonary artery. The data developed from the dual frequency measurements allow the absolute heart output to be determined, thereby accounting for the variability of fluid capacity of mixing volume.
Two other prior art works are relevant to the present invention. In his M.S. thesis entitled, "Electronic Augmentation of Thermo-dilution Techniques," Massachusetts Institute of Technology, Cambridge, September 1975, L. M. Rubin discloses a technique for processing thermo-dilution data wherein the assumes that the input heating signal and output blood temperature signal are exactly in phase and makes no corrections for phase shifting due to mixing volume or drift (of the baseline blood temperature). In U.S. Pat. No. 4,507,974, M. L. Yelderman applied a pseudo-random heating signal to the blood and reconstructed a dilution curve similar to that of the regular thermo-dilution technique, by cross correlating the temperature and power signals. The patient's cardiac output was then obtained by Yelderman by integrating the area under the curve, just as is done in a conventional thermo-dilution measurement.
Neither Webler, Yelderman, nor Newbower et al. teach how to minimize a potentially significant source of error that can occur during injectateless cardiac output measurements--namely the error caused by slow baseline blood temperature fluctuations. Newbower et al. does not mention the problem, and Webler's measurement of cardiac output specifically assumes that the baseline blood temperature remains constant, so that the change in the temperature of blood in the pulmonary artery is entirely due to effect of the input signal. However, the baseline blood temperature may be changing during the measurement of cardiac output, so as to cause the value for the temperature, T, of the blood leaving the heart to be in error. Such changes in the baseline temperature of the blood may be due to a long term periodic cycling of the patient's body temperature or may simply result from the patient's body warming from a chilled condition. Certain medical procedures include chilling a patient before surgery to slow metabolism, and in this instance, the change in baseline blood temperature may be attributable to a gradual warming of the patient to normal body temperature (37.degree. C.). In either case, whether due to a long term periodic cycle or a gradual linear drift, the change in baseline blood temperature approximates a ramp wave form. This gradual variation in blood temperature can contribute a significant error to the measurement of cardiac output.
It is therefore preferable that a non-injectate method for determining cardiac output be compensated for a drift or slow fluctuations (noise) in the baseline temperature of blood entering the heart. Compensation for the effect of drift in the baseline temperature while determining cardiac output should preferably also account for a phase lag or time delay in the determination of the cardiac output, which is introduced by the mixing volume of the heart and for other time delays that arise in the measurement.